1. Field of the Invention
The present invention in general concerns magnetic resonance tomography (MRT) as employed in medicine for examination of patients. The present invention in particular concerns an MRT method with over-sampling in at least one phase coding direction of an image region to be visualized.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear magnetic resonance and has been successfully used as an imaging method for over 15 years in medicine and biophysics. In this modality, the subject is exposed to a strong, constant magnetic field. The nuclear spins of the atoms in the subject, which were previously randomly oriented, thereby align. Radio-frequency energy can now excite these ordered nuclear spins to a specific oscillation. This oscillation generates the actual measurement signal, which is acquired by appropriate reception coils. By the use of non-homogeneous magnetic fields generated by gradient coils, the measurement subject can be spatially coded in all three spatial directions, generally known as “spatial coding”.
The acquisition of the data ensues in MRT in k-space (frequency domain). The MRT image in the image domain is linked with the MRT data in k-space by means of Fourier transformation. The spatial coding of the subject, which spans k-space, ensues in all three spatial directions by means of gradients. Differentiation is made between the slice selection (establishes an acquisition slice in the subject, typically the z-axis), the frequency coding (establishes a direction in the slice, typically the x-axis) and the phase coding (determines the second dimension within the slice, typically the y-axis). Moreover, the selected slice can be sub-divided into further slices by an additional second phase coding along the z-axis.
A slice is thus first selectively excited (for example in the z-direction) and a phase coding in the z-direction is possibly conducted. The coding of the spatial information in the slice ensues with a combined phase and frequency coding by means of both of these aforementioned orthogonal gradient fields, that in the example of a slice excited in the z-direction are generated by the gradient coils in the x-direction and the y-direction, respectively.
A possible pulse sequence to acquire the data in an MRT experiment is shown in FIGS. 2A and 2B. The sequence is a spin-echo sequence. In this, the magnetization of the spin is tilted in the x-y plane by a 90° excitation pulse. In the course of time (½ TE; TE is the echo time) a dephasing occurs of the magnetization components that mutually form the transverse magnetization in the x-y plane Mxy. After a certain time (for example, ½ TE), a 180° pulse is radiated in the x-y plane such that the dephased magnetization components are reflected without the precession direction and precession speed of the individual magnetization portions being changed. After a further time duration TD, which may be ½ TE, the magnetization components again point in the same direction, i.e. a regeneration of the transverse magnetization occurs, designated as a “rephasing”. The complete regeneration of the transverse magnetization is designated as a spin-echo.
In order to measure an entire slice of the subject to be examined, the imaging sequence is repeated N-times for different values of the phase coding gradient, for example Gy. The temporal separation of the respectively excited RF pulses is designated as a repetition time TR. The magnetic resonance signal (spin-echo signal) is likewise sampled, digitized, and stored N times in every sequence repetition via the Δt-clocked ADO (analog-digital converter) in equidistant time steps Δt in the presence of the read-out gradient Gx. In this manner, a numerical matrix is created row by row (matrix in k-space, or k-matrix) with N×N data points, as shown in FIG. 2B. An MR image of the slice in question with a resolution of N×N pixels can be directly reconstructed from this data set via a Fourier transformation (a symmetric matrix with N×N points is only one example, asymmetrical matrices can be generated as well).
The measured MRT signal, which defines a value of the k-matrix is described by the amplitude, the frequency and the phase of the signal. In MR tomography, the amplitude contains the information about the spin density, while frequency and phase of the signal are used for spatial coding of the respective spatial directions.
In contrast to the frequency coding, the phase coding gradient (for example Gy is only between excitation and acquisition for a fixed duration. All spins do in fact precess again with the same resonance frequency after the switching of the phase coding gradient, but they now possess a spatially-dependent phase. As can be seen in FIG. 2A, the duration of the phase coding gradient is kept constant so that the phase depends only on the spatial direction (for example y-direction) to be coded as well as on the respective gradient amplitude Gy. The phase of the signal changes linearly with the spatial direction (y-direction). A defined phase modulation of the nuclear magnetic resonance signals thus ensues for every gradient amplitude.
So that the association of the phase modulation is unambiguous for each phase coding step, and thus for the entirety of all phase coding steps, the region of interest to be measured of the subject to be examined cannot exceed a delimited range in the phase coding direction. The range in which signals can be unambiguously associated with a position is designated as a field of view (FOV). When the region of interest (ROI) or the subject itself lies within the FOV, all subject positions are unambiguously determined via the phase modulation of a plurality of phase coding steps. If the region of interest lies partially outside of the FOV, it leads to ambiguities. In the image the portion of the ROI or the subject that is beyond an edge of the FOV is projected over the other side of the portion within the FOV. This is known as a foldover artifact (also called aliasing or backfolding or phase wrapping or wrap around artifacts).
Foldovers thus arise in subject structures that lie in the measurement slice but outside of the image field (FOV; normally rectangular) marked by the user in the measurement slice in the phase coding direction. This is due to the fact that the phase coding gradient exhibits a periodicity and is only distinct from 0° to 360°. This fact is illustrated using FIGS. 3A and 3B: the MRT apparatus cannot differentiate between 370° and 10°, which is why the part 33 of the subject 32 (which, for example, protrudes to the right from the image region (FOV) 31 in FIG. 3A in the phase coding direction at 370°) is folded (wrapped around) into the image region again at 10° on the left side 33′ of the reconstructed image 31′ that contains the subject 32′, as shown in FIG. 3B. In reverse, the part 34 of the subject 32 (that, for example, likewise protrudes to the left from the image region 31 in FIG. 3A in the phase coding direction at −5°) folds into the image region 31′ (at 355°) again on the right side 34′ of the reconstructed image 31′.
The subject structure 32′ completely located in the FOV, however, is unambiguously imaged in the reconstructed image.
The simplest manner to prevent or avoid these foldovers is to align the phase coding direction such that subject structures no longer protrude into this from the image region. In the most common case—see FIG. 4A, in which the image region (the FOV) is surrounded on all sides by subject structures—this is not possible.
In this case, a method according to the prior art proceeds as follows:
After an overview slice image of the entire subject in the measurement plane of interest has been acquired in an initial “scout image” procedure, and the image region of interest FOVa of the width “a” has been indicated by the user in the phase coding direction, both sides of the initial FOVa are expanded from its center point so far that a Fovb old is created which completely contains the subject.
In terms of calculation, the expansion in the phase coding direction bold is determined from the overall width of the subject in the phase coding direction s as well as the distance d of the center point of the starting image region FOVa to the center point of the subject in the phase coding direction according to the equation
                              b          old                =                              s            2                    +                      s            2                    +                      2            ⁢                                        d                                                                            =                  s          +                      2            ⁢                                        d                                                        
In order to not commit to a specific phase coding direction, an image enlargement in the frequency coding directly (vertical in FIGS. 4A, 4B) is likewise effected according to the same method (i.e., replacing s with z and replacing d with the vertical offset of the center of FOVa from the center of the subject region 32), such that ultimately the image region is expanded so far that a subject region 32 no longer protrudes from the new FOVb old. This has the consequence that foldovers are not generated anymore given a scan of the widened FOVb old parallel or perpendicular to a in the initial image region FOVa.
A significant disadvantage of this method is the fact that an oversampling of the current image region of interest FOVa of width “a” occurs (4B), that not only lengthens the scan duration by multiple times, but also demands multiple times the memory storage space.